The invention relates generally to nuclear magnetic resonance (NMR) and magnetic resonance imaging (MRI), and more particularly to NMR and MRI at ultralow magnetic fields.
Nuclear magnetic resonance (NMR) is a technique for obtaining information about atoms and the molecules they form. NMR operates on atoms having nuclei in which at least one proton or neutron is unpaired. This imbalance causes these nuclei to spin on an axis like miniature tops and gives rise to a magnetic moment, i.e. the nuclei behave like magnets with north and south poles.
When exposed to an external magnetic field, these spinning magnets attempt to align their axes along the lines of magnetic force. The alignment is not exact, however, resulting in a wobbly rotation (precession) about the force lines that is unique for each type of nuclei. If, while exposed to the magnetic field, the nuclei are bombarded with radio (RF) waves, they will absorb and re-emit energy at a specific frequency according to their rate of rotation. This resonating frequency therefore becomes a signature signal by which the nuclei can be identified.
When nuclei absorb the energy of an incoming radio wave, they are knocked out of alignment with the external magnetic field lines. As they subsequently lose this energy, the nuclei come back into alignment. The rate at which resonating nuclei realign themselves with magnetic field lines provides detailed information on their position and motion with respect to neighboring nuclei. This provides a noninvasive technique to study the structural, dynamic, and spatial relationships of atoms in a sample of molecules.
NMR has two basic subsets—spectroscopy and imaging. In NMR spectroscopy, the frequency of the incoming radio wave is varied, and all of the different frequencies absorbed and emitted by the nuclei are measured to obtain a resonance spectrum. This NMR spectrum reveals the molecular makeup of the material down to the respective positions and motions of the constituent atoms.
In magnetic resonance imaging (MRI), the frequency of the incoming radio wave is kept constant, but the strength of the external magnetic field is varied. The resulting signal corresponds to the total number of spinning nuclei present in any part of the sample, i.e. the atomic density of the sample at that point. Information obtained from an array of points can be translated by computer into a recognizable image.
Since the invention of MRI in the early 1970s, MRI scanners have steadily developed towards higher magnetic field strengths. The enhanced sensitivity attainable at high field makes it possible to resolve features at ever shorter length scales, and enables fast imaging experiments with close to real-time resolution. State-of-the-art clinical scanners operate at a field strength of 1.5 T, corresponding to a proton Larmor frequency of 64 MHz; currently, there is a drive to gain approval for 4 T imagers for clinical use. A number of facilities around the world now have 7 T scanners for research purposes.
At the same time, the last three decades have seen continued effort toward the development of systems for MRI in low magnetic fields. Much of this work has been motivated by considerations of cost: a commercial full-body imager operating at 1.5 T costs several million dollars, and the operation of such a machine places considerable demands on the infrastructure of the hospital or research facility. In addition, due to the size and complexity of the high-field system, it must necessarily remain fixed in one location, and the sample or subject must be transported to the system and inserted into the confining bore of the high field magnet; in certain cases this is simply not possible. A low-cost, portable MRI scanner is extremely appealing, as is an open MRI system, which would enable acquisition of MRIs at the same time that a medical procedure is performed. Inexpensive, portable imagers would enable MRI to address a wide variety of new problems, potentially transforming it from a highly specialized clinical and research technique to a much more widespread, flexible tool for rapid patient screening and general noninvasive imaging. However, any sort of portable or open MRI system would need to operate at relatively low magnetic field strengths.
Moreover, despite the serious disadvantage of reduced sensitivity, the images acquired in low field should, in principle, be of higher quality than those acquired in high magnetic field. An inevitable drawback of high-field imaging is that of susceptibility artifacts. When a heterogeneous sample is placed in a magnetic field, variations in magnetic susceptibility over the sample volume give rise to spurious magnetic field gradients. When these spurious gradients become comparable to the gradients that are used for encoding, the image is severely distorted. In medical imaging, the presence of dental fillings or jewelry is enough to destroy the MRI; abrupt changes in susceptibility at solid-liquid and solid-air interfaces inside the body, such as in the sinuses, produce distortions which are more subtle, but which nevertheless place strict limits on the achievable spatial resolution. Since the strength of the spurious gradients scales linearly with the strength of the applied field, it is possible to eliminate susceptibility-induced distortions entirely by imaging in low magnetic field.
Finally, T1 contrast in tissue is enhanced in low magnetic field. Because of this, low-field images allow sharper differentiation of different organs and tissue types, and potentially contain richer information than the corresponding images acquired in high field. (It is interesting to note that, in the early days of MRI, many researchers were skeptical that high-field MRI would ever amount to a useful clinical tool, precisely because of the degradation of tissue contrast in high field.)
There have been a number approaches to low-field MRI in recent years. These have generally relied on Faraday detection in a static field of order 10 mT to 100 mT, which is generated by an electromagnet. The main obstacle in these studies is the low sensitivity intrinsic to the low field experiment. In a different approach, H. C. Seton et al., “A 4.2 K receiver coil and SQUID amplifier used to improve the SNR of low-field magnetic resonance images of the human arm,” Meas. Sci. Technol. 8, 198–207 (1997) employed a tuned SQUID magnetometer for NMR detection; the SQUID provided an SNR enhancement of a factor of 2.8–4.5 over conventional detection in images acquired from room temperature samples in a field of 10 mT. In the low-field imaging work of A. Macovski et al., “Novel approaches to low cost MRI,” Magn. Reson. Med. 30, 221–230 (1993) and W. Shao et al., “Low readout field magnetic resonance imaging of hyperpolarized xenon and water in a single system,” Appl. Phys. Lett. 80, 2032–2034 (2002), spins were prepolarized in a field of 0.3 T, while the NMR signals were detected in a much lower field of 30 mT. Here, the homogeneity of the polarizing field was not crucial, and the prepolarization step led to an enhancement of sample magnetization by an order of magnitude. Using similar techniques, J. Stepi{hacek over (s)}nik et al., “NMR imaging in the Earth's magnetic field,” Magn. Reson. Med. 15, 386–391 (1990) and G. Planinsic et al., “Relaxation-time measurement and imaging in the Earth's magnetic field,” J. Magn. Reson. Ser. A 110, 170–174 (1994), acquired MRIs in the magnetic field of the Earth (BEarth˜50 μT), demonstrating the enhanced T1 contrast attainable in low-field. In both the works of Macovski et al. and Stepi{hacek over (s)}nik et al., however, Faraday detection in the field of the Earth entailed substantial signal loss.
Superconducting Quantum Interference Devices (SQUIDs) are sensitive detectors of magnetic fields based on the quantum mechanical Josephson effect. SQUIDs are based on superconductors, whose resistance drops to zero when cooled to a critical temperature Tc. A SQUID is formed by separating its superconducting material with a very thin insulating barrier through which electron pairs can tunnel. This combination of superconducting material and insulating barrier forms a Josephson junction, i.e. two superconductors joined by a weak link. The SQUID consists of a superconducting ring or square interrupted in two spots by Josephson junctions. When sufficient electrical current is applied to the SQUID, a voltage is generated across its body. In the presence of a magnetic field, this voltage will change as the strength of the field changes. Thus the SQUID turns a change in a magnetic field, which is more difficult to measure, into a change in voltage, which is very easy to measure.
For application purposes, SQUIDs are almost always coupled to auxiliary components. To form a magnetometer, a SQUID is connected to a flux transformer, a device consisting of a relatively large loop of superconducting material and a much smaller multiturn coil. Since the large loop picks up a magnetic field over a much greater area, the sensitivity of the SQUID to changes in magnetic field strength is boosted manyfold.
Originally SQUIDs were made with low Tc superconductors, e.g. niobium (Tc=9.5K), which required cooling with liquid helium. More recently, high Tc SQUIDs have been made, using high Tc ceramic oxide superconducting materials, e.g. yttrium barium copper oxide (YBCO) materials (Tc=93K), which only require cooling with liquid nitrogen, which is much less expensive and easier to work with than liquid helium. A high Tc low noise SQUID is described in U.S. Pat. No. 6,023,161 issued Feb. 8, 2000.
Low transition temperature SQUIDs have been used experimentally to detect NMR and nuclear quadrupole resonance (NQR) signals, e.g. Dinh M. Ton That et al., “Direct current superconducting quantum interference device spectrometer for pulsed nuclear magnetic resonance and nuclear quadrupole resonance at frequencies up to 5 MHz,” Rev. Sci. Instr. 67, 2890 (1996). Low Tc SQUIDs have been used to image polarized helium and xenon at relatively low fields, e.g. M. P. Augustine et al., “Low field magnetic resonance images of polarized noble gases obtained with a dc superconducting quantum interference device,” Appl. Phys. Lett. 72 (15), 1908 (1998). The feasibility of using a high Tc SQUID to detect NMR signals has been demonstrated, S. Kumar et al., “Nuclear magnetic resonance using a high temperature superconducting quantum interference device,” Appl. Phys. Lett. 70 (8), 1037 (1997).
SQUIDs were first used in the 1980s to detect NMR signals in low magnetic field. However, the majority of SQUID NMR studies have been performed on samples in the solid state, at liquid helium (LHe) temperatures. Recently, there has been increased interest in extending SQUID NMR techniques to samples in the liquid state, and in particular to systems which are biologically relevant. S. Kumar et al., “Broadband SQUID NMR with room temperature samples,” J. Magn. Reson. B 107, 252 (1995) demonstrated NMR spectra from animal tissue measured at room temperature. H. C. Seton et al., ibid., used SQUIDs to image room temperature samples in a field of 10 mT, and K. Schlenga et al., “Low-field magnetic resonance imaging with a high-Tc dc superconducting quantum interference device,” Appl. Phys. Lett. 75, 3695–3697 (1999) and U.S. Pat. No. 6,159,444 issued Dec. 12, 2000, used a SQUID magnetometer fabricated from the high transition temperature superconductor YBCO to image thermally polarized proton samples at room temperature in a field of 2 mT. Despite these early efforts, however, SQUID NMR studies of liquid samples remain extremely limited in number and in scope. The central challenge for SQUID NMR studies of liquids is that of low sensitivity. Thermal polarizations are two orders of magnitude lower at 300 K than at 4.2 K. Moreover, in order to cool the SQUID below its superconducting transition temperature, it is necessary to thermally isolate the detector from the sample; filling factor is therefore quite low.
The NMR effect is produced by a spin magnetic moment on nuclei in a sample. A magnetic field causes the spin magnetic moments to precess around the field at the Larmor frequency ω which is proportional to the magnetic field.
In low field NMR (typically ≦10 mT) the spin precesses at correspondingly low frequencies, typically below 500 kHz, around the field direction. In conventional NMR, in which a resonant circuit is used to detect the precessing magnetization, the induced voltage signal V is proportional to the spin magnetization M and its rate of change (frequency) ω. Since M is also proportional to the frequency ω, V scales with ω2. As a result it is difficult to detect NMR signals at low fields with a conventional Faraday detector. In contrast, SQUIDs can be used to measure magnetic flux directly, resulting in much higher signal to noise (S/N) ratio at low frequencies. However, the use of SQUIDs for NMR/MRI has heretofore been limited, and not used at ultralow magnetic fields of tens or hundreds of microtesla.